MRI is used in radiology to visualize details of structures in a patient's body. For aligning the magnetic spin of the nuclei, mostly of protons in water molecules in the body tissue, the patient is placed inside a powerful static magnetic field. Excited by an electro-magnetic radio-frequency pulse from a transmitter coil the nuclei resonating at this frequency deflect and then gradually relax towards the static field while emitting detectable electro-magnetic radiation, which can be captured as an “echo” at a certain time after excitation (the “echo time”) by a receiver coil. Relaxation times and the resonance frequency of the nuclei depend both on local properties of the tissue material, which represents the underlying principle allowing visualization of these properties. The characteristics of the image are also influenced by proton density and magnetic field strength. By superposing lateral and longitudinal gradient fields and adjusting the excitation frequency, certain volumetric regions (“volume elements”) can be measured by both selective excitation and frequency analysis of the captured echo signals. Fourier-transforming the latter generates complex representations of the frequency values and their phases, which can be made readable.
When the excitation is performed on 2-dimensional slices through the object, they are selected by adjusting the gradients in the magnetic field strength, thereafter resonant frequency values are captured and processed, forming 2-dimensional images, a number of which can be merged to form a 3-dimensional representation of the object. Alternatively, in 3-dimensional imaging, a large volume of tissue is excited and spatially encoded using frequency encoding in one direction and phase encoding in both of the remaining two orthogonal directions.
In the past, magnitude values have been used primarily from the complex representations. Nevertheless, phase values allow for the extraction of additional information about local properties inside the tissue, where they specifically benefit from strong susceptibility effects at high magnetic field strengths. For example, phase information is used in neuroimaging in phase-contrast angiography, Susceptibility-Weighted Imaging (SWI), susceptibility mapping—also known as Quantitative Susceptibility Mapping (QSM), Susceptibility Tensor Imaging, to depict iron accumulation in neurodegenerative disorders and to map in vivo conductivity. It can also be used to monitor temperature and encode flow velocity.
However, each phase value acquired by a receiver coil of the MRI machine is subject to a time-independent offset, often referred to as “phase offset”. The phase offset comprises spatially constant components, e.g. due to the cable length from a receiver coil to a receiver, as well as spatially variable components, e.g. due to the path lengths of the excitation and echo signals from particular locations in the object to the receiver coil in question, reflecting inhomogeneities in the magnetic field.
It has been an aim of research to eliminate the effects of the phase offset, e.g. in order to facilitate combining multiple phase images acquired with a plurality of receiver coils arranged in an array around the object, and thereby increase the quality of an acquired image in terms of its signal-to-noise ratio. Several approaches have been presented in the past: One of these approaches, presented by Hammond, K. E. et al., “Development of a robust method for generating 7.0 T multichannel phase images of the brain with application to normal volunteers and patients with neurological diseases”, NeuroImage 2008, 39; pp. 1682-1692, suggests to estimate a spatially constant phase offset by setting the phase values to zero in all coils at the centre of an image. This method, while being easy to apply, results in areas of poor phase matching. An alternative solution is to refer the phase values of each receiver coil to a “virtual reference coil” which is the result of a two-step procedure. In the first step, a combined image (the Virtual Reference Coil, or VRC, image) is generated using an image-based constant (as in the method of Hammond et al.). In the second step, the phase image from each coil is referenced to the VRC image. While the matching of phase values of different receiver coils is very good in this case, local field variations arising from magnetic susceptibility are not reflected. Yet another solution, proposed by Roemer, P. B. et al., “The NMR phased array”, Magn Reson Med 1990, 16; pp. 192-225, uses an additional body coil or other homogeneous volume reference coil—i.e. a coil which is separate from said receiver coils and has to be sensitive over (at least) all the tissue over which the receiver coils, arranged in an array around the object, are sensitive—for referencing and using a phase offset separately measured by means of the body coil for each receiver coil; however, such an additional body coil is not commonly available in ultra-high field scanners and requires extra space and control.